Direct conversion energy discriminating ct detector

ABSTRACT

A CT detector capable of energy discrimination and direct conversion is disclosed. The detector includes multiple layers of semiconductor material with the layers having varying thicknesses. The detector is constructed to be segmented in the x-ray penetration direction so as to optimize count rate performance as well as avoid saturation.

CROSS REFERENCE TO RELATED APPLICATIONS

The present application is continuation of and claims priority of U.S.Ser. No. 10/848,877 filed May 19, 2004, the disclosure of which isincorporated herein by reference.

BACKGROUND OF THE INVENTION

The present invention relates generally to diagnostic imaging and, moreparticularly, to a multi-layer direct conversion CT detector capable ofproviding photon count and/or energy data with improved saturationcharacteristics.

Typically, in radiographic imaging systems, an x-ray source emits x-raystoward a subject or object, such as a patient or a piece of luggage.Hereinafter, the terms “subject” and “object” may be interchangeablyused to describe anything capable of being imaged. The beam, after beingattenuated by the subject, impinges upon an array of radiationdetectors. The intensity of the attenuated beam radiation received atthe detector array is typically dependent upon the attenuation of thex-rays. Each detector element of the detector array produces a separateelectrical signal indicative of the attenuated beam received by eachdetector element. The electrical signals are transmitted to a dataprocessing system for analysis which ultimately produces an image.

In some computed tomography (CT) imaging systems, the x-ray source andthe detector array are rotated about a gantry within an imaging planeand around the subject. X-ray sources typically include x-ray tubes,which emit the x-rays as a beam at a focal point. X-ray detectorstypically include a collimator for collimating x-ray beams received atthe detector, a scintillator for converting x-rays to light energyadjacent the collimator, and a photodiode for receiving the light energyfrom an adjacent scintillator and producing electrical signalstherefrom. Typically, each scintillator of a scintillator array convertsx-rays to light energy. Each photodiode detects the light energy andgenerates a corresponding electrical signal. The outputs of thephotodiodes are then transmitted to the data processing system for imagereconstruction.

Conventional CT imaging systems utilize detectors that convertradiographic energy into current signals that are integrated over a timeperiod, then measured and ultimately digitized. A drawback of suchdetectors is their inability to provide data or feedback as to thenumber and/or energy of photons detected. That is, conventional CTdetectors have a scintillator component and photodiode component whereinthe scintillator component illuminates upon reception of radiographicenergy and the photodiode detects illumination of the scintillatorcomponent and provides an electrical signal as a function of theintensity of illumination. While it is generally recognized that CTimaging would not be a viable diagnostic imaging tool without theadvancements achieved with conventional CT detector design, a drawbackof these detectors is their inability to provide energy discriminatorydata or otherwise count the number and/or measure the energy of photonsactually received by a given detector element or pixel. Accordingly,recent detector developments have included the design of an energydiscriminating, direct conversion detector that can provide photoncounting and/or energy discriminating feedback. In this regard, thedetector can be caused to operate in an x-ray counting mode, an energymeasurement mode of each x-ray event, or both.

These energy discriminating, direct conversion detectors are capable ofnot only x-ray counting, but also providing a measurement of the energylevel of each x-ray detected. While a number of materials may be used inthe construction of a direct conversion energy discriminating detector,semiconductors have been shown to be one preferred material. A drawbackof direct conversion semiconductor detectors, however, is that thesetypes of detectors cannot count at the very high x-ray photon flux ratestypically encountered with conventional CT systems. The very high x-rayphoton flux rate ultimately leads to detector saturation. That is, thesedetectors typically saturate at relatively low x-ray flux levels. Thissaturation can occur at detector locations wherein small subjectthickness is interposed between the detector and the radiographic energysource or x-ray tube. It has been shown that these saturated regionscorrespond to paths of low subject thickness near or outside the widthof the subject projected onto the detector fan-arc. In many instances,the subject is more or less circular or elliptical in the effect onattenuation of the x-ray flux and subsequent incident intensity to thedetector. In this case, the saturated regions represent two disjointedregions at extremes of the fan-arc. In other less typical, but not rareinstances, saturation occurs at other locations and in more than twodisjointed regions of the detector. In the case of an ellipticalsubject, the saturation at the edges of the fan-arc is reduced by theimposition of a bowtie filter between the subject and the x-ray source.Typically, the filter is constructed to match the shape of the subjectin such a way as to equalize total attenuation, filter and subject,across the fan-arc. The flux incident to the detector is then relativelyuniform across the fan-arc and does not result in saturation. What canbe problematic, however, is that the bowtie filter may not be optimumgiven that a subject population is significantly less than uniform andnot exactly elliptical in shape. In such cases, it is possible for oneor more disjointed regions of saturation to occur or conversely toover-filter the x-ray flux and create regions of very low flux. Lowx-ray flux in the projection will ultimately contribute to noise in thereconstructed image of the subject.

A number of imaging techniques have been developed to address saturationof any part of the detector. These techniques include maintenance of lowx-ray flux across the width of a detector array, for example, by usinglow tube current or current that is modulated per view. However, thissolution leads to increased scanned time. That is, there is a penaltythat the acquisition time for the image is increased in proportion tothe nominal flux needed to acquire a certain number of x-rays that meetimage quality requirements. Other solutions include the implementationof an over-range algorithm that is used to generate replacement data forthe saturated data. However, these algorithms may imperfectly replacethe saturated data as well as contribute to the complexity of the CTsystem.

It would therefore be desirable to design a direct conversion, energydiscriminating CT detector that does not saturate at the x-ray photonflux rates typically found in conventional CT systems.

BRIEF DESCRIPTION OF THE INVENTION

The present invention is directed to a multilayer CT detector that canbe made to perform at very high count rates that overcomes theaforementioned drawbacks.

A CT detector capable of energy discrimination and direct conversion isdisclosed. The detector includes multiple layers of semiconductormaterial of varying thicknesses throughout the detector. In this regard,the detector is constructed so as to be segmented in the x-raypenetration direction to optimize count rate performance as well asavoid saturation.

The CT detector supports not only x-ray photon counting, but energymeasurement or tagging as well. As a result, the present inventionsupports the acquisition of both anatomical detail as well as tissuecharacterization information. In this regard, the energy discriminatoryinformation or data may be used to reduce the effects of beam hardeningand the like. Further, these detectors support the acquisition of tissuediscriminatory data and therefore provide diagnostic information that isindicative of disease or other pathologies. For example, detection ofcalcium in a plaque in a view is possible. This detector can also beused to detect, measure, and characterize materials that may be injectedinto a subject such as contrast agents and other specialized materialssuch as targeting agents. Contrast materials can, for example, includeiodine that is injected into the blood stream for better visualization.A method of fabricating such a detector is also disclosed.

Therefore, in accordance with one aspect of the present invention, adirect conversion CT detector includes multiple direct conversion layersdesigned to directly convert radiographic energy to electrical signalsrepresentative of energy sensitive CT data. The detector also includesan electrical signal collection layer sandwiched between adjacent directconversion layers.

In accordance with another aspect, the present invention includes a CTsystem having a rotatable gantry having a bore centrally disposedtherein, a table movable fore and aft through the bore and configured toposition a subject for CT data acquisition, and a radiographic energyprojection source positioned within the rotatable gantry and configuredto project radiographic energy toward the subject. The CT system alsoincludes a detector array disposed within the rotatable gantry andconfigured to detect radiographic energy projected by the projectionsource and impinged by the subject. The detector array includes aplurality of detector cells, wherein each cell has a stacked arrangementof semiconductor layers in a direction generally that of energyprojection and designed to provide energy sensitive data acquired fromthe subject in response to receiving radiographic energy.

According to another aspect, the present invention includes a CTdetector having a first means and a second means for directly convertingradiographic energy to electrical signals. The detector also has meansfor receiving electrical signals interstitially positioned between thefirst means for directly converting and the second means for directlyconverting.

Various other features, objects and advantages of the present inventionwill be made apparent from the following detailed description and thedrawings.

BRIEF DESCRIPTION OF THE DRAWINGS

The drawings illustrate one preferred embodiment presently contemplatedfor carrying out the invention.

In the drawings:

FIG. 1 is a pictorial view of a CT imaging system.

FIG. 2 is a block schematic diagram of the system illustrated in FIG. 1.

FIG. 3 is a perspective view of one embodiment of a CT system detectorassembly.

FIG. 4 is a perspective view of a CT detector.

FIG. 5 is illustrative of various configurations of the detector in FIG.4 in a four-slice mode.

FIG. 6 is a partial perspective view of a two-layer director inaccordance with the present invention.

FIG. 7 is a cross-sectional view of FIG. 6 taken along lines 7-7thereof.

FIGS. 8-10 illustrate cross-sectional views of direct conversiondetectors in accordance with several additional embodiments of thepresent invention.

FIG. 11 is a cross-sectional schematic view of that shown in FIG. 10illustrating signal feedthroughs that are created in another embodimentof the invention.

FIG. 12 is a pictorial view of a CT system for use with a non-invasivepackage inspection system.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

The operating environment of the present invention is described withrespect to a four-slice computed tomography (CT) system. However, itwill be appreciated by those skilled in the art that the presentinvention is equally applicable for use with single-slice or othermulti-slice configurations. Moreover, the present invention will bedescribed with respect to the detection and conversion of x-rays.However, one skilled in the art will further appreciate that the presentinvention is equally applicable for the detection and conversion ofother radiographic energy.

Referring to FIGS. 1 and 2, a computed tomography (CT) imaging system 10is shown as including a gantry 12 representative of a “third generation”CT scanner. Gantry 12 has an x-ray source 14 that projects a beam ofx-rays 16 toward a detector assembly 18 on the opposite side of thegantry 12. Detector assembly 18 is formed by a plurality of detectors 20which together sense the projected x-rays that pass through a medicalpatient 22. Each detector 20 produces an electrical signal thatrepresents not only the intensity of an impinging x-ray beam but is alsocapable of providing photon or x-ray count data, and hence theattenuated beam as it passes through the patient 22. During a scan toacquire x-ray projection data, gantry 12 and the components mountedthereon rotate about a center of rotation 24.

Rotation of gantry 12 and the operation of x-ray source 14 are governedby a control mechanism 26 of CT system 10. Control mechanism 26 includesan x-ray controller 28 that provides power and timing signals to anx-ray source 14 and a gantry motor controller 30 that controls therotational speed and position of gantry 12. A data acquisition system(DAS) 32 in control mechanism 26 review data from detectors 20 andconverts the data to digital signals for subsequent processing. An imagereconstructor 34 receives sampled and digitized x-ray data from DAS 32and performs high speed reconstruction. The reconstructed image isapplied as an input to a computer 36 which stores the image in a massstorage device 38.

Computer 36 also receives commands and scanning parameters from anoperator via console 40 that has a keyboard. An associated cathode raytube display 42 allows the operator to observe the reconstructed imageand other data from computer 36. The operator supplied commands andparameters are used by computer 36 to provide control signals andinformation to DAS 32, x-ray controller 28 and gantry motor controller30.

In addition, computer 36 operates a table motor controller 44 whichcontrols a motorized table 46 to position patient 22 and gantry 12.Particularly, table 46 moves portions of patient 22 through a gantryopening 48.

As shown in FIGS. 3 and 4, detector assembly 18 includes a plurality ofdetectors 20, with each detector including a number of detector elements50 arranged in a cellular array. A collimator (not shown) is positionedto collimate x-rays 16 before such beams impinge upon the detectorassembly 18. In one embodiment, shown in FIG. 3, detector assembly 18includes 57 detectors 20, each detector 20 having an array size of16×16. As a result, assembly 18 has 16 rows and 912 columns (16×57detectors) which allows 16 simultaneous slices of data to be collectedwith each rotation of gantry 12.

Switch arrays 54 and 56, FIG. 4, are multi-dimensional semiconductorarrays coupled between cellular array 52 and DAS 32. Switch arrays 54and 56 include a plurality of field effect transistors (FET) (not shown)arranged as multi-dimensional array and are designed to combine theoutputs of multiple cells to minimize the number of data acquisitionchannels and associated cost. The FET array includes a number ofelectrical leads connected to each of the respective detector elements50 and a number of output leads electrically connected to DAS 32 via aflexible electrical interface 58. Particularly, about one-half ofdetector element outputs are electrically connected to switch 54 withthe other one-half of detector element outputs electrically connected toswitch 56. Each detector 20 is secured to a detector frame 60, FIG. 3,by mounting brackets 62.

It is contemplated and recognized that for some applications, the countrate limitation of the FET arrays may make them less desirable. In thisregard, as will be described, each detection pixel or cell is connectedto a channel of electronics.

Switch arrays 80 and 82 further include a decoder (not shown) thatenables, disables, or combines detector element outputs in accordancewith a desired number of slices and slice resolutions for each slice.Decoder, in one embodiment, is a decoder chip or a FET controller asknown in the art. Decoder includes a plurality of output and controllines coupled to switch arrays 54 and 56 and DAS 32. In one embodimentdefined as a 16 slice mode, decoder enables switch arrays 54 and 56 sothat all rows of the detector assembly 18 are activated, resulting in 16simultaneous slices of data for processing by DAS 32. Of course, manyother slice combinations are possible. For example, decoder may alsoselect from other slice modes, including one, two, and four-slice modes.

As shown in FIG. 5, by transmitting the appropriate decoderinstructions, switch arrays 54 and 56 can be configured in thefour-slice mode so that the data is collected from four slices of one ormore rows of detector assembly 18. Depending upon the specificconfiguration of switch arrays 54 and 56, various combinations ofdetectors 20 can be enabled, disabled, or combined so that the slicethickness may consist of one, two, three, or four rows of detectorelements 50. Additional examples include, a single slice mode includingone slice with slices ranging from 1.25 mm thick to 20 mm thick, and atwo slice mode including two slices with slices ranging from 1.25 mmthick to 10 mm thick. Additional modes beyond those described arecontemplated.

As described above, each detector 20 is designed to directly convertradiographic energy to electrical signals containing energydiscriminatory data. The present invention contemplates a number ofconfigurations for these detectors. Notwithstanding the distinctionsbetween each of these embodiments, each detector does share two commonfeatures. One of these features is the multilayer arrangement ofsemiconductor films or layers. In a preferred embodiment, eachsemiconductor film is fabricated from Cadmium Zinc Telluride (CZT).However, one skilled in the art will readily recognize that othermaterials capable of the direct conversion of radiographic energy may beused. The other common feature between the various embodiments is theuse of interstitial or intervening metallized films or layers separatingthe semi-conducting layers. As will be described, these metallizedlayers are used to apply a voltage across a semiconductor layer as wellas collect electrical signals from a semiconductor layer.

It is generally well known that photon count rate performance of asemiconductor is a function of the square of the thickness of thedetector and the radiographic energy deposition process is exponential.The count rate performance for a CZT detector may be defined by:$T_{TR} = {\frac{L^{2}}{V\quad\mu_{e}}.}$

From this definition, assuming a thickness of L=0.3 cm and an electricfield V of 1000 V/cm, and with a μ_(e) of about 1000, a maximum countrate of 1.0 megacounts may be achieved. In other words, the count rateof a CZT semiconductor layer that is 3 mm thick may have a count rateperformance in the range of 1.0 megacounts/sec. However, as will bedescribed, constructing a direct conversion semiconductor detector withmultiple layers as opposed to a single thicker layer can improve countrate performance.

Referring now to FIG. 6, a portion of a two-layered CZT or directconversion detector 20 a in accordance with one embodiment of thepresent invention is shown in perspective. Detector 20 a is defined by afirst semiconductor layer 62 and a second semiconductor layer 64. Duringthe fabrication process, each semiconductor layer 62, 64 is constructedto have a number of electronically pixilated structures or pixels todefine a number of detection elements 65. This electronic pixilation isaccomplished by applying a 2D array 67, 69 of electrical contacts 65onto a layer 62, 64 of direct conversion material. Moreover, in apreferred embodiment, this pixilation is defined two-dimensionallyacross the width and length of each semiconductor layer 62, 64.

Detector 20 a includes a contiguous high voltage electrode 66, 68 forsemiconductor layers 62, 64, respectively. Each high voltage electrode66, 68 is connected to a power supply (not shown) and is designed topower a respective semiconductor layer during the x-ray or gamma raydetection process. One skilled in the art will appreciate that each highvoltage connection layer should be relatively thin so as to reduce thex-ray absorption characteristics of each layer and, in a preferredembodiment, is a few hundred angstroms thick. As will be described ingreater detail below, these high voltage electrodes may be affixed to asemiconductor layer through a metallization process.

Referring now to FIG. 7, a cross-sectional view of FIG. 6 taken alongline 7-7 thereof illustrates the relative thickness of eachsemiconductor layer 62, 64. Similar to the high voltage electrode layers66, 68, the 2D arrays 67, 69 should also be minimally absorbent ofradiographic energy. Each array or signal collection layer is designedto provide a mechanism for outputting the electrical signals created bythe semiconductor layers to a data acquisition system or other systemelectronics. One skilled in the art will appreciate that several hundredinterconnects (not shown) are used to connect each contact 65 with theCT system electronics.

In addition, as shown in FIG. 7, the thickness of the semiconductorlayers 62, 64 is different from one another. In this regard, more x-raysare deposited in semiconductor layer 62 than in semiconductor layer 64.For example, assuming that semiconductor layer 62 has a thickness of onemillimeter (mm) and semiconductor 64 has a thickness of 2 mm,semiconductor layer 62 is expected to absorb about 78% of the x-rayswhereas the second semiconductor layer 64 is expected to absorb about22% of the x-rays. Further, it is expected that the first semiconductorlayer 62 is to experience a maximum count rate that is approximatelynine times faster than that of a single layer semiconductor 3 mm thick.However, the first semiconductor layer 62 measures only approximately78% of the total flux thereby providing an 11.5 times increase ineffective max count rate performance compared to a single semiconductorlayer 3 mm thick. The second semiconductor layer 64 is expected to havea count rate that is 2.25 times faster than that of a single 3 mm thicksemiconductor but measures only approximately 22% of the total flux,thereby, providing an equivalent or effective max count rate that isapproximately 10.2 times that expected to be experienced with a singlelayer of semiconductor material 3 mm thick. As a result of the improvedcount rates of the segmented detector described above relative to asingle layer of semiconductor material, detector 20 a may be constructedto provide a tenfold increase in count rate performance.

The above dimensions are illustrative of the improvement in maximumcount rate that may be experienced with a two layer detector. However,it is contemplated that more than two layers may be used to construct aCT detector with improved count rate characteristics. For example, asingle 0.43 mm layer is expected to absorb about 54% of x-rays receivedand, as such, has a maximum count rate of approximately 40 times that ofa single layer, 3.0 mm thick semiconductor. However, a 0.43 mm layerabsorbs only approximately 54% of the total flux to provide anequivalent or effective max count rate of approximately 92 times that ofa single semiconductor layer that is 3 mm thick. Additional layers maybe added to provide an overall count rate increase of 9200%.

Referring now to FIG. 8, another contemplated design for a CZT or directconversion detector is shown. In this embodiment, detector 20 b alsoincludes a pair of semiconductor layers 74, 76. In contrast to thepreviously described embodiment, detector 20 b includes a single, commonsignal collection layer or 2D contact array 78. This single, yet commonarray 78 is designed to collect electrical signals from bothsemiconductor layers 74, 76 and output those electrical signals to a DASor other system electronics. In addition, detector 20 b includes a pairof high voltage electrodes 80, 82. Each high voltage electrodeeffectively operates as a cathode whereas the contacts of the 2D array78 operate as an anode. In this regard, the voltage applied via highvoltage connections 80, 82 creates a circuit through each semiconductorlayer to the signal collection contacts array 78.

Yet another contemplated embodiment is illustrated in FIG. 9. As shownin this embodiment, detector 20 c includes four semiconductor layers 84,86, 88, and 90. Detector 20 c further includes two electricallyconductive lines or paths 92, 94 that are electrically connected to highvoltage electrodes 87, 89, 91 as well as collection contact arrays 93,95. Electrically conductive path 92 receives and translates electricalsignals from contact arrays 93, 95. In this regard, a single data outputis provided to the CT system's electronics. Similar to a single signalcollection lead, a single high voltage connection 94 is used to powerthe four semiconductor layers 84-90 via electrodes 87, 89, 91. Detector20 c only requires a single high voltage connection.

Referring to FIG. 10, a monolithic embodiment of the present inventionis shown. Similar to the embodiment of FIG. 7, detector 20 d includesfour semiconductor layers 96-102. Each semiconductor layer 96-102 isconnected to a pair of electrically conductive layers. In this regard,one electrically conductive layer is used for application of a voltagewhereas the other electrically conductive layer is used for collectionof the electrical signals generated by the respective semiconductorlayers. To minimize the number of electrically conductive layers,detector 20 d utilizes an alternating electrically conductive layerarchitecture. That is, every other electrically conductive layer is usedfor high voltage connection with the other electrically conductivelayers used for signal collection. In this regard, electricallyconductive layers 104, 106, and 108 are used for application of arelatively high voltage whereas layers 110 and 112 include contacts forsignal collection. As such, high voltage collection layers 104 and 108are used to apply a voltage to semiconductor layers 96 and 102,respectively. High voltage connection layer 106 is used to apply avoltage to semiconductor layers 98 and 100.

As described above, in a preferred embodiment, each semiconductor layeris constructed from CZT material. One skilled in the art will appreciatethat there are a number of techniques that may be used to construct sucha semiconductor. For example, molecular beam epitaxy (MBE) is one methodthat may be used to grow each thin layer of CZT material. One skilled inthe art will appreciate that a number of techniques may be used tometallize the semiconductor layers to provide the electricallyconductive connections heretofore described.

Further, metallization may also be used to provide signal feedthroughsfor the collection contacts as illustrated in FIG. 11. As shown, asingle layer of semiconductor material 114 is sandwiched between anarray 116 of collection contacts and a high voltage electrode layer 118.Prior to metallization of the semiconductor layer 114 to form collectioncontact array 116 and high voltage electrode layer 118, holes 120 may beetched or otherwise formed in semiconductor 114. The holes 120 may thenbe metallized to provide a signal feed path 122 from a respectivecollection contact 124. The signal feedthroughs or conductive paths 122are constructed within holes 120 so as to not be in contact with thenear-contiguous high voltage electrode layer 118. In this regard, signalruns may extend vertically or in the x-ray reception directionthroughout the detector to a bus (not shown) designed to translate theelectrical signals emitted by the individual collection contacts 124 tothe CT system's electronics. As a result, a stacked arrangement of aseries of thin stacked layers in the x-ray direction is formed.

Referring now to FIG. 12, package/baggage inspection system 126 includesa rotatable gantry 128 having an opening 130 therein through whichpackages or pieces of baggage may pass. The rotatable gantry 128 housesa high frequency electromagnetic energy source 132 as well as a detectorassembly 134. A conveyor system 136 is also provided and includes aconveyor belt 138 supported by structure 140 to automatically andcontinuously pass packages or baggage pieces 142 through opening 130 tobe scanned. Objects 142 are fed through opening 130 by conveyor belt138, imaging data is then acquired, and the conveyor belt 138 removesthe packages 142 from opening 130 in a controlled and continuous manner.As a result, postal inspectors, baggage handlers, and other securitypersonnel may non-invasively inspect the contents of packages 142 forexplosives, knives, guns, contraband, etc.

Therefore, a direct conversion CT detector includes multiple directconversion layers designed to directly convert radiographic energy toelectrical signals representative of energy sensitive CT data. Thedetector also includes an electrical signal collection layer sandwichedbetween adjacent direct conversion layers.

The present invention also includes a CT system having a rotatablegantry having a bore centrally disposed therein, a table movable foreand aft through the bore and configured to position a subject for CTdata acquisition, and a radiographic energy projection source positionedwithin the rotatable gantry and configured to project radiographicenergy toward the subject. The CT system also includes a detector arraydisposed within the rotatable gantry and configured to detectradiographic energy projected by the projection source and impinged bythe subject. The detector array includes a plurality of detector cells,wherein each cell has a stacked arrangement of semiconductor layers in adirection generally that of energy projection and designed to provideenergy sensitive data acquired from the subject in response to receivingradiographic energy.

The present invention further includes a CT detector having a firstmeans and a second means for directly converting radiographic energy toelectrical signals. The detector also has means for receiving electricalsignals interstitially positioned between the first means for directlyconverting and the second means for directly converting.

The present invention has been described in terms of the preferredembodiment, and it is recognized that equivalents, alternatives, andmodifications, aside from those expressly stated, are possible andwithin the scope of the appending claims.

1. A CT detector comprising: a first semiconductor layer spaced apartfrom a second semiconductor layer, the semiconductor layers designed toconvert x-rays to electrical signals and collectively having a maxphoton count rate that is greater than a single semiconductor layerhaving a thickness equal to a composite thickness of the first and thesecond semiconductor layers; a readout layer sandwiched between thefirst and the second semiconductor layers, the readout layer designed totranslate the electrical signals to a data acquisition system.
 2. The CTdetector of claim 1 further comprising a voltage application layerpositioned adjacent to a surface of each semiconductor layer opposite tothat which the readout layer is positioned.
 3. The CT detector of claim2 wherein each semiconductor layer is in contact with a readout layerand a voltage application layer.
 4. The CT detector of claim 1 whereinthe semiconductor layers are formed of at least CZT.
 5. The CT detectorof claim 1 wherein the semiconductor layers are arranged in a layeredstack in an x-ray penetration direction.
 6. The CT detector of claim 5wherein the first semiconductor layer absorbs more x-rays than thesecond semiconductor layer which is positioned farther from an x-raysource.
 7. The CT detector of claim 1 wherein the semiconductor layershave the same thickness.
 8. The CT detector of claim 1 wherein the firstsemiconductor layer is thicker than the second semiconductor layer. 9.The CT detector of claim 1 wherein the readout layer reads outelectrical signals from both semiconductor layers.
 10. A CT detectorcomprising: a stacked arrangement of direct conversion layers aligned todirectly convert x-rays to electrical signals; the stacked arrangementincluding a first direct conversion layer and a second direct conversionlayer spaced apart from one another by a shared readout layer; the firstdirect conversion layer more absorbent of x-rays than the second directconversion layer.
 11. The CT drector of claim 10 wherein the readoutlayer carries energy-sensitive CT data from the first and the seconddirect conversion layers to a data acquisition system.
 12. The CTdetector of claim 11 wherein the energy-sensitive CT data includesphoton count and energy level information.
 13. The CT detector of claim10 wherein the stacked arrangement further includes a pair of biasinglayers, each biasing layer connected to a respective direct conversionlayer and designed to create a potential across a direct conversionlayer with the shared readout layer.
 14. The CT detector of claim 13wherein the readout layer and the biasing layers are metallized to arespective surface of the first and the second direct conversion layers.